Cardiovascular implants of enhanced biocompatibility

ABSTRACT

Cardiovascular and other medical implants fabricated from low-modulus Ti--Nb--Zr alloys to provide enhanced biocompatibility and hemocompatibility. The cardiovascular implants may be surface hardened by oxygen or nitrogen diffusion or by coating with a tightly adherent, hard, wear-resistant, hemocompatible ceramic coating. The cardiovascular implants include heart valves, total artificial heart implants, ventricular assist devices, vascular grafts, stents, electrical signal carrying devices such as pacemaker and neurological leads, defibrillator leads, and the like. It is contemplated that the Ti--Nb--Zr alloy can be substituted as a fabrication material for any cardiovascular implant that either comes into contact with blood thereby demanding high levels of hemocompatibility, or that is subject to microfretting, corrosion, or other wear and so that a low modulus metal with a corrosion-resistant, hardened surface would be desirable.

RELATED APPLICATIONS

This is a division of application Ser. No. 08/112,599 filed on Aug. 26,1993, issued as U.S. Pat. No. 5,477,864 which is a continuation-in-partof U.S. Ser. No. 08/036,414 filed on Mar. 24, 1993, issued U.S. Pat. No.5,509,933 which is in turn a continuation-in-part of U.S. Ser. No.07/986,280, filed on Dec. 7, 1992, abandoned, which is acontinuation-in-part of Ser. No. 07/647,453, filed on Jan. 28, 1991,issued as U.S. Pat. No. 5,169,597, which is in turn a continuation ofU.S. Ser. No. 07/454,181, filed on Dec. 21, 1989, abandoned.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to a range of cardiovascular and other implantsfabricated of metallic alloys of enhanced hemocompatibility that canoptionally be surface hardened to provide resistance to wear, orcold-worked or cold-drawn to reduce elastic modulus, if necessary. Morespecifically, the invention is of synthetic heart valves, ventricularassist devices, total artificial hearts, stents, grafts, pacers,pacemaker leads and other electrical leads and sensors, defibrillators,guide wires and catheters, and percutaneous devices fabricated ofTi--Nb--Zr alloys.

2. Description of the Related Art

Cardiovascular implants have unique blood biocompatibility requirementsto ensure that the device is not rejected (as in the case of naturaltissue materials for heart valves and grafts for heart transplants) orthat adverse thrombogenic (clotting) or hemodynamic (blood flow)responses are avoided.

Cardiovascular implants, such as heart valves, can be fabricated fromnatural tissue. These bioprostheses can be affected by gradualcalcification leading to the eventual stiffening and tearing of theimplant.

Non-bioprosthetic implants are fabricated from materials such aspyrolytic carbon-coated graphite, pyrolytic carbon-coated titanium,stainless steel, cobalt-chrome alloys, cobalt-nickel alloys, aluminacoated with polypropylene and poly-4-fluoroethylene.

For synthetic mechanical cardiovascular devices, properties such as thesurface finish, flow characteristics, surface structure, charge, wear,and mechanical integrity all play a role in the ultimate success of thedevice. For example, typical materials used for balls and discs forheart valves include nylon, silicone, hollow titanium, TEFLON™,polyacetal, graphite, and pyrolytic carbon. Artificial hearts andventricular assist devices are fabricated from various combinations ofstainless steel, cobalt alloy, titanium, Ti--6Al--4V alloy, carbon fiberreinforced composites, polyurethanes, BIOLON™ (DuPont), HEMOTHANE™(Sarns/3M), DACRON™, polysulfone, and other thermoplastics. Pacers,defibrillators, leads, and other similar cardiovascular implants aremade of Ni--Co--Cr alloy, Co--Cr--Mo alloy, titanium, and Ti--6Al--4Valloy, stainless steel, and various biocompatible polymers. Stents andvascular grafts are often made of DACRON™ stainless steel or otherpolymers. Catheters and guide wires are constructed from Co--Ni orstainless steel wire with surrounding polymer walls.

One of the most significant problems encountered in heart valves,artificial hearts, assist devices, pacers, leads, stents, and othercardiovascular implants is the formation of blood clots(thrombogenesis). Protein coatings are sometimes employed to reduce therisk of blood clot formation. Heparin is also used as ananti-thrombogenic coating.

It has been found that stagnant flow regions in the devices ornon-optimal materials contribute to the formation of blood clots. Thesestagnant regions can be minimized by optimizing surface smoothness andminimizing abrupt changes in the size of the cross section through whichthe blood flows or minimizing other flow interference aspects. Whilematerials selection for synthetic heart valves, and cardiovascularimplants generally, is therefore dictated by a requirement for bloodcompatibility to avoid the formation of blood clots (thrombus),cardiovascular implants must also be designed to optimize blood flow andwear resistance.

Even beyond the limitations on materials imposed by the requirements ofblood biocompatibility and limitations to designs imposed by the need tooptimize blood flow, there is a need for durable designs since it ishighly desirable to avoid the risk of a second surgical procedure toimplant cardiovascular devices. Further, a catastrophic failure of animplanted device will almost certainly result in the death of thepatient.

The most popular current heart valve designs include the St. Judemedical tilting disc double cusp (hi-leaf) valve. This valve includes acircular ring-like pyrolytic carbon valve housing or frame and a flowcontrol element which includes pyrolytic carbon half-discs or leavesthat pivot inside the housing to open and close the valve. The twoleaves have a low profile and open to 85° from the horizontal axis.

Another popular heart valve is the Medtronic-Hall Valve wherein the flowcontrol element is a single tilting disc made of carbon coated withpyrolytic carbon which pivots over a central strut inside a solidtitanium ring-like housing. A third, less popular design, is theOmniscience valve Which has a single pyrolytic disc as a flow controlelement inside a titanium housing. Finally, the Starr-Edwards ball andcage valves have a silastic ball riding inside a cobalt-chrome alloycage. The cage is affixed to one side of a ring-like body for attachmentto the heart tissue. More recent designs include trileaflet designs andconcave bileaflet designs to improve blood flow.

From the point of view of durability, heart valves made oflow-thrombogenic pyrolyte carbon could fail from disc or pivot jointwear or fracture related to uneven pyrolytic carbon coating, fracture ofthe ball cage, disc impingement, strut wear, disc wear, hinge failure,and weld failure. A more recent heart valve, the Baruah Bileaflet issimilar to the St. Jude design but opens to 80° and is made of zirconiummetal. The valve has worked well over its approximately two-year historywith roughly 200 implants to date in India. This performance can bepartly attributed to the lower elastic modulus of zirconium( about 90GPa) and the resultant lower contact stress severity factor (Cc of about0.28×10⁻⁷ m) when the disc contacts the frame. In contrast, pyrolyticconstructions produce contact stress severity factors of about 0.54×10⁻⁷m.

Although zirconium has worked well to date and can reduce contact stressseverity, zirconium metal is relatively soft and sensitive to frettingwear. This is partly due to hard, loosely attached, naturally-presentpassive oxide surface films (several nanometers in thickness) which caninitiate microabrasion and wear of the softer underlying metal. However,this naturally present zirconium oxide passive film is thrombogenicallycompatible with blood and the design is acceptable from a hemodynamicstandpoint. Therefore, while the zirconium bileaflet valve appears tomeet at least two of the major requirements for cardiac valve implants,namely blood compatibility and design for minimum stagnant flow regions,the use of soft zirconium metal leads to a relatively high rate offretting wear and leads to the expectation that the valve may be lessdurable than one produced from materials less susceptible to frettingwear. Titanium and titanium alloys present a similar limitation, andCo--Cr--Mo, stainless steel, and Co--Ni alloys have much greater elasticmodulus.

There exists a need for a metallic cardiac valve implant that isbiocompatible, compatible with blood in that it does not induce bloodclotting and does not form a calcified scale, that is designed tominimize stagnant flow areas where blood clotting can be initiated, thathas a low elastic modulus for lower contact stress severity factors toensure resistance to wear from impact, and that has a surface that isalso resistant to microabrasion thereby enhancing durability.

Heart diseases, many of which cannot be cured by conventional surgery ordrug therapy, continue to be a leading cause of death. For the seriouslyill patient, heart replacement is often one of the few viable optionsavailable.

In 1988, the NHBLI began funding research and development forpermanently implantable, electrically driven, total artificial hearts(TAHs). The pumping mechanism of the TAHs would be implanted into thechest cavity of the patient and the device would be powered by a batterypack and a small transformer, worn by the patient, which transmitsenergy to the heart with no physical connections through the skin. TheNHBLI funded four separate groups: The Cleveland Clinic Foundation &Nimbus, Inc.; Pennsylvania State University & Sarns/3M; The Texas HeartInstitute and Abiomed, Inc.; and the University of Utah. Consequently,four competing designs were developed.

The development of TAHs posed several issues. Firstly, it was necessaryto duplicate the action of a human heart, ensure long-term reliabilityand biocompatibility, while producing a device that fits into the chestcavity in terms of both its total volume and the orientation of itsconnections to natural vessels in the body. Aside from the purelymechanical, wear, and power supply issues, it is also necessary that thedesign and materials prevent infection and thrombosis. Blood is anon-Newtonian fluid and its properties, such as viscosity, change withoxygen content, kidney function, and even the age of the patient.Further, plasma contains a suspension of fragile red blood cells whichmay be caught in artificial valves, or other mechanically stressfulareas, thereby destroying these cells. It is therefore necessary todevelop a TAH that does not stress blood components, and to fabricatethe pump from materials that are not only biocompatible, but also "bloodcompatible" in the sense of minimizing damage to blood components andminimizing the formation of blood clots.

Many of the above comments also apply to ventricular assist devices(VADs), one of which is being developed by the Novacor Division ofBaxter Health Care Corp. In the use of a VAD, the patient's heartremains in place while the VAD boosts the pumping pressure of the leftventricle of the heart. Consequently, the VAD is an assist device ratherthan a replacement. However, the VAD must be blood compatible for thesame reasons as the total artificial heart.

There exists a need for a material that is lightweight, readily formableinto complex shapes, biocompatible, and blood and tissue compatible witha hard surface that is resistant to abrasive wear, microfretting wear,and the corrosive effects of body fluids, for use in heart assist orreplacement devices (including EMHs, VADs, and TAHs) to prolong the lifeof mechanical components while at the same time minimizing anydeleterious effect on blood components.

SUMMARY OF THE INVENTION

The invention provides cardiovascular and other medical implants of alow modulus, biocompatible, hemocompatible, metallic alloy of titaniumwith niobium and optionally zirconium. The invention implants includeheart valves, artificial hearts, ventricular assist devices,defibrillators, pacers, electrical leads, sensors, grafts, stents, andcatheter devices. The invention also provides surface hardened versionsof these devices produced by oxygen or nitrogen diffusion hardening toimprove resistance to cavitation, microfretting wear, and impact-inducedwear.

The inherently low modulus of Ti--Nb--Zr alloys, between about 6 toabout 12 million psi depending on metallurgical treatment andcomposition, provide a more flexible and forgiving construct forcardiovascular applications while improving contact stress levels, valveclosure, and the ability of leaves in certain valve designs toself-align with blood flow and reduce thrombodynamic effects.

The invention provides components for use in mechanical heartreplacement or assist devices, such as external mechanical hearts(EMHs), total artificial hearts (TAHs), and ventricular assist devices(VADs), that are lightweight, while also being resistant to corrosivebody fluids, mechanical wear, abrasive wear, and microfretting wear.Further, the components have much improved blood compatibility in thesense of reduced risk of thrombogenesis (blood clotting).

The preferred low modulus titanium alloys of the invention have thecompositions: (i) titanium; about 10 wt. % to about 20 wt. % niobium;and optionally from about 0 wt. % to about 20 wt. % zirconium; and (ii)titanium; about 35 wt. % to about 50 wt. % niobium; and optionally fromabout 0 wt. % to about 20 wt. % zirconium. Tantalum can also be presentas a substitute for Nb. These alloys are referred to herein as"Ti--Nb--Zr alloys," even though tantalum may also be present.

The exclusion of elements besides titanium, zirconium, and niobium, ortantalum results in an alloy which does not contain known toxins orcarcinogens, or elements that are known or suspected of inducingdiseases or adverse tissue response in the long term.

Without the presence of zirconium in the composition, the ability of theTi--Nb--Zr alloy to surface harden during oxygen or nitrogen diffusionhardening treatments is more limited. Therefore, presence of zirconiumis especially preferred when the alloy implant must be diffusionhardened. Other non-toxic filler materials such as tantalum, whichstabilize the β-phase of titanium alloy, but do not affect the lowmodulus, i.e., maintain it at less than about 85 GPa, could also beadded.

A porous coating, such as plasma-sprayed or sintered titanium ortitanium alloy (including Ti--Nb--Zr alloy) beads or wire mesh may alsobe added to the implant's surfaces to improve tissue attachment, such asthe formation of an endothelial cell layer, preferred in artificialheart, ventricular assist devices, grafts, and stent devices. Suchcoatings provide more favorable blood interaction and flowcharacteristics, and also tend to stabilize the implant with the body.Thus, such porous coatings may also be useful for connecting regions ofthese devices as well as for heart valves and grafts. Even though theapplication of such porous coatings usually requires sintering atrelatively high temperatures, the properties of the Ti--Nb--Zr alloythat might affect its usefulness as an implant material are notadversely affected.

BRIEF DESCRIPTION OF THE DRAWINGS

A better understanding of the present invention can be obtained when thefollowing detailed description of the preferred embodiments isconsidered in conjunction with the following drawings, in which:

FIG. 1 shows a simplified representation of a ball valve, like theStarr-Edwards Valve.

FIG. 2 is a simplified representation of a disc valve.

FIG. 3 is a simplified representation of a tilting disc, single cuspvalve like the Medtronic-Hall valve.

FIG. 4 is a simplified representation of a tilting disc, double cusp orbileaflet valve of the St. Jude or Baruah type.

FIG. 5 is a schematic diagram of the Penn State/Sarns/3M design totalartificial heart.

FIG. 6 is a schematic diagram of the University of Utah total artificialheart.

FIG. 7 is a schematic diagram of the Cleveland Clinic/Nimbus, Inc. totalartificial heart.

FIG. 8A is a schematic diagram, in cross section, of the Texas HeartInstitute/Abiomed total artificial heart.

FIG. 8B is a side view, in cross section, of the Texas HeartInstitute/Abiomed total artificial heart of FIG. 8A.

FIGS. 9A and B are schematic diagrams of end and front views,respectively, of a ventricular assist device.

FIG. 10 is a schematic diagram of a vascular graft of woven metallicwire composition.

FIG. 11A is a schematic diagram showing a balloon expandable stentpositioned within a segment of a blood vessel to be propped open.

FIG. 11B is a schematic diagram showing a balloon expanding a stent intoposition within a blood vessel.

FIG. 11C is a schematic diagram of a stent expanded in a blood vessel.

FIG. 12A is a schematic diagram of the components of a defibrillator,showing power source, lead wire, and polymeric patch with coiledelectrode.

FIG. 12B is a cross section of the lead wire of FIG. 12A.

FIG. 13A is a schematic diagram, in partial cross section, of the distalend of a guide wire.

FIG. 13B is a cross-section of the guide wire with coating thicknessexaggerated.

FIG. 13C shows a catheter containing coiled wire and polymer wall.

FIGS. 14A-C are schematic diagrams of prior art pacemaker leads withpolyurethane covering.

FIG. 14D is a schematic of an embodiment of the invention Ti--Nb--Zrpacemaker leads.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

The implants of the invention are fabricated from an alloy containingtitanium as a component. The preferred low modulus titanium alloys havethe compositions: (i) titanium, about 10 wt. % to about 20 wt. %niobium, and optionally from about 0 wt. % to about 20 wt. % zirconium;and (ii) titanium, about 35 wt. % to about 50 wt. % niobium, andoptionally from about 0 wt. % to about 20 wt. % zirconium.

In a preferred embodiment wherein the implants are surface hardened byoxygen or nitrogen diffusion, zirconium is beneficially present inamounts ranging from about 0.5 to about 20 wt. %.

Even though it is apparent that the titanium proportion of alloy used tomake the invention implants could be less than 50 wt. %, nevertheless,for the purposes of this specification, it is referred to as a"Ti--Nb--Zr alloy" or a "titanium alloy." The alloy most preferablycomprises about 74 wt. % titanium in combination with about 13 wt. % ofzirconium and 13 wt. % of niobium. While tantalum may be substituted forniobium to stabilize β-phase titanium, niobium is the preferredcomponent due to its effect of lowering the elastic modulus of the alloywhen present in certain specific proportions. Other elements are notdeliberately added to the alloy but may be present in such quantitiesthat occur as impurities in the commercially pure titanium, zirconium,niobium, or tantalum used to prepare the alloy and such contaminants asmay arise from the alloying process.

In the specification-and claims, the term "high strength" refers to analloy having a tensile strength above at least about 620 MPa.

The term "low modulus," as used in the specification and claims, refersto a Young's modulus below about 90 GPa.

Although the hot rolled, reheated, and quenched Ti--Nb--Zr alloy is asuitable implant material, its properties can be improved by forging orother metallurgical processes or an aging heat treatment or acombination of these. Aging treatment can increase the strength andhardness of the material, and reduce its elongation while maintaining arelatively low modulus of elasticity. The treatment can be varied toobtain the desired properties. U.S. Pat. No. 5,169,597 to Davidson, etal., hereby fully incorporated by reference, deals in more detail withthe useful Ti--Nb--Zr alloys. Further, U.S. Ser. No. 08/036,414, herebyfully incorporated by reference, teaches how to hot work Ti--Nb--Zralloys to produce high strength, low modulus medical implants.

It may be desirable for other reasons, such as reducing microfrettingwear between mating mechanical components, to surface harden the alloyimplants using oxygen or nitrogen diffusion hardening methods, orcoating with a hard wear resistant coating. In the latter event, thesurface of the prosthesis may be coated with an amorphous diamond-likecarbon coating or ceramic-like coating such as zirconium or titaniumoxide, zirconium or titanium nitride, or zirconium or titanium carbideusing chemical or plasma vapor deposition techniques to provide a hard,impervious, smooth surface coating. These coatings are especially usefulif the prosthesis is subjected to conditions of wear, such as, forinstance, in the case of mating parts in artificial hearts orventricular assist devices, or as an electrically insulative coating onelectrical leads (i.e., pacemaker, defibrillator, neurological, sensors)of Ti--Nb--Zr alloy.

Methods for providing hard, low-friction, impervious, biocompatibleamorphous diamond-like carbon coatings are known in the art and aredisclosed in, for example, EPO patent application 302 717 A1 to Ion Techand Chemical Abstract 43655P, Vol. 101, describing Japan Kokai 59/851 toSumitomo Electric, all of which are incorporated by reference herein asthough full set forth.

A preferred process for oxygen diffusion hardening is described in ourco-pending patent application U.S. Ser. No. 08/112,612 (temporarilyidentified by Attorney Docket No. A93096US), now U.S. Pat. No. 5,372,660filed on the same day as this patent application, which is hereby fullyincorporated by reference. Oxygen diffusion hardening according to thisprocess requires the supply of oxygen, or an oxygen containingatmosphere, or compounds partially composed of oxygen, such as water(steam), carbon dioxide, nitrogen dioxide, sulfur dioxide, and the like.These substances are supplied to the implant to be hardened which ismaintained at a temperature preferably between 200° C. and 1200° C. Theamount of time required at a given temperature to effectively producethe surface and near-surface hardened implants is related exponentially,by an Arhennius-type relationship to the temperature. That is, shorterperiods of time-are required at higher temperatures for effectivediffusion hardening. The resultant oxygen diffusion hardened implantsare characterized in that the oxide film contains primarily a mixture oftitanium and zirconium oxides in the implant surface. Niobium oxides mayalso be present. Immediately underlying this mixed-oxide film issometimes a region of oxygen-rich metal alloy. Underlying thesometimes-obtained oxygen-rich alloy layer is the core Ti--Ni--Zr alloy.The interface between the sometimes-obtained oxygen-rich alloy layer andthe oxide regions is typically zirconium-rich in comparison to theunderlying Ti--Ni--Zr alloy. In a most preferred embodiment, theTi--Ni--Zr alloy is subjected to temperature and an environment of argongas that has been moisturized by bubbling through a water bath. Thewater vapor disassociates at the implant surface to produce oxygen whichdiffuses into the implant to produce the desired hardened surface.

Nitrogen diffusion processes can also be utilized in which nitrogensources are provided instead of oxygen. These nitrogen diffusion surfacehardening processes will tend to harden the metal alloy substrate in asimilar manner to that of oxygen diffusion hardening or conventionaloxygen hardening (which is also useful), and produce a yellow-orangeinsulative, wear-resistant surface oxide instead of the blue-blacksurface oxide which typically forms from the in situ oxygen diffusionhardening process.

Implants fabricated from the inventive alloy may be supplied with aporous bead or wire coating of titanium alloy of the same or differentcomposition including pure titanium to allow endothelial cell attachmentto blood-contacting flow surfaces or for stabilization of the implant inthe body of the patient after implantation by tissue ingrowth into theporous structure. Such porous structures are normally attached to theimplant surface by sintering or plasma spraying. Sintering involvesheating the implant to above about 1250° C. The mechanical properties oftitanium alloys can change significantly due to substantial grain growthand other metallurgical factors arising from the sintering process.Thus, after sintering to attach the porous coating, it may be desirablein some instances to reheat the Ti--Nb--Zr implant, for example, toabout 875° C. (or above the β-transus) for 20-40 minutes then waterquench before aging at about 500° C. for about 6 hours to restoremechanical properties. If quenched adequately from the sinteringtemperature, it may be possible to go directly to the aging process. Analternative method of attaching a porous coating is to simply plasmaspray metal powder or micro-beads onto the implant's surface afterappropriate mechanical and thermal treatments.

Further, the implants of the invention may optionally be surface coatedwith medicaments such as anti-inflammatory agents, anti-thrombus agents,antibiotics, proteins that reduce platelet adhesion, and the like toimprove their acceptability in a living body.

Heart Valves

In its simplest form, a synthetic cardiac valve includes a valve bodyfor affixing the valve to the body tissue and through which blood flows,and a flow control element for allowing or blocking off blood flow. Forinstance, FIG. 4 shows a typical bileaflet valve having a valve bodythat includes a ring-like housing 400 with an inner ring 402 that hastwo flanges 404 each containing two slots for receiving hinges attachedto leaflets. The flow control element of this valve comprises twoleaflets 405, in the approximate shape of half discs, with hingeelements attached at diametrically opposite ends. These hinge elementsfit within apertures or slots in the flanges 404 of the inner ring 402and are able to rotate through less than 180° C. in these apertures.Thus, in operation, the flow control elements are in the position shownin FIG. 4 with the valve open with blood flowing from top to bottom.When blood flow reverses and flows from bottom to top, the bileaflets405 pivot about their hinges to close the apertures in the ring-likevalve body. Consequently, there is a significant amount of movementabout the hinge elements and slots where microfretting wear might beinitiated. Furthermore, the bileaflet half disc flow control elements405 may impinge upon the inner ring 402 of the valve body, therebyleading to cavitation or impact-induced wear.

The invention provides heart valves of various designs, exemplified inFIG. 1-4, each comprising parts subject to impact and wear that arefabricated from Ti--Nb--Zr alloy. More recent designs (not shown in theFigures) include concave bileaflet and trileaflet designs which areintended to improve blood flow. The heart valves are preferablysubjected to a hardening process, such as oxygen or nitrogen diffusionhardening to produce a harder Ti--Nb--Zr surface that is resistant tomicrofretting wear at hinge points and impact wear at those locationswhere a flow control element impacts the valve body. Consequently, theinvention valves have a longer cycle life than the currently used St.Jude, Omniscience, Starn-Edwards, Medtronic-Hall, or Baruah valves.Indeed, as mentioned before, the Baruah valve is currently fabricated ofzirconium or zirconium alloys and would therefore be subject torelatively rapid wear because of the relative softness of zirconium andits alloys. The use of Ti--Nb--Zr alloy compositions also provides alow-thrombus, blood-compatible surface favorable in reducing theincidence of blood clots.

Artificial Hearts/Ventricular Assist Devices

FIG. 5 is illustrative of the Penn State design which incorporates astainless-steel roller screw 10 positioned between two flexiblediaphragm blood pumps (right side pump 12 is shown, the other is withinshell 8). A high speed, low torque brushless DC motor causes the rollerscrew 10 to turn thereby moving the roller screw shafts 16 linearly backand forth. To each end of the guide shaft 16 is attached a pusher plate18. When the pusher plate 18 in the right side pump moves backward,blood is drawn into the pump space 14. At the same time, the pusherplate in the corresponding left side pump moves towards the left pushingblood out of its pump space. In this type of pump, the pusher plates 18act against a flexible membrane 12, which is in contact with blood, andwhich can be inflated on the pump suction stroke and deflated on pumpdischarge stroke. The pusher plates are driven by the roller screw 10and, thus, according to the invention, six revolutions of roller screw10 are required for a full stroke with a motor speed of about 3,000 RPM.While planetary rollers are inserted between a roller screw nut androller screw 10 to give rolling, not sliding contact and to spreadmechanical load over many contact points, and while the entire rollerscrew system is hardened to minimize wear, the reliability of thissystem is improved by the use of the invention components. Thus, rollerscrew 10 and the roller-screw nut with which it is in rolling contactare both fabricated of Ti--Nb--Zr alloy and the surfaces are hardened orcoated with a hard, tightly adherent coating. Further, the surfaces ofthe pump nozzles 2 shown as 4 and 6, connecting elements 20 and conduits22, into and from which blood is continuously being pumped, isfabricated from Ti--Nb--Zr to reduce adverse reaction with blood tissueon those surfaces presented to blood components to minimize thepotential for the formation of thrombus and blood clots.

FIG. 6 is a schematic cross-sectional diagram of the University of Utahelectrohydraulic heart. This heart includes shells 50 and 52 with a pumpmotor 34 interposed between them. In this type of heart, a motor 20 isused to pressurize silicone oil in regions 22 and 24 on the undersidesof flexible diagrams 26 and 28, respectively, to move blood in and outof the chambers 30 and 32 above the flexible diaphragms. For example,when motor 34 pressurizes silicone oil into chamber 22, then flexiblediaphragm 26 expands upward and outwardly to push blood flow out ofchamber 30 in direction 36. At the same time, silicon oil flows out ofchamber 24 towards chamber 22 thereby allowing flexible diaphragm 28 toassume a natural position, shown in FIG. 2, and drawing blood intochamber 32 as shown from direction 38. Upon reversal of the direction ofthe bidirectional pump 34, the opposite effects are achieved.

Since the University of Utah pump is of a bidirectional design, andtypically operates at speeds between 10,000-13,000 RPM in the highpressure direction and 5,000-8,000 RPM in the reverse, moving componentsof the pump are subject to microfretting and mechanical wear. Therefore,the invention components for the bidirectional axial flow pump 34 usedin the University of Utah TAH design are fabricated from Ti--Nb--Zralloy coated hardened or with a wear resistant, hard coating that istightly adherent, to reduce wear of the high speed components. Further,surfaces 40, 41, 42 43, 44 and 45 are in direct contact with blood andare made of Ti--Nb--Zr alloy to improve blood compatibility and reducethe potential for thrombus and blood clotting. Thus, the shells of theheart 50 and 52 are also fabricated of Ti--Nb--Zr alloy to reducethrombogenesis.

FIG. 7 is a schematic cross-sectional illustration of the ClevelandClinic TAH which utilizes a motor to turn a gear pump 56 which provideshydraulic pressure at about 100 psi to cause reciprocal movement ofactuators 58 which in turn drive pusher-plates 60 that act on flexiblediaphragms 62 to pump the blood. The actuators 58 operate slidinglywithin guide sleeve 64 so that wear on contact surfaces between actuatorand sleeve may be expected. Further, the TAH has a flow reversing valve64 with machine elements, such as bearing surfaces, subject to wear whenthe TAH is in use. Thus, TAH elements that are subject to wear and thatmay be advantageously fabricated of Ti--Nb--Zr alloys that are thensurface hardened and/or coated with a hard, wear resistant, tightlyadherent coating, include the guide sleeve 64, the actuators 58, thepump's gear elements and shaft and the rotary valve 64. Moreover, toreduce the risk of erosion damage to the pump from cavitation, the pumphousing 72 may likewise be fabricated of Ti--Nb--Zr alloys. Finally,internal surfaces 66, 68 of the heart housing 70 are in direct bloodcontact. Thus it is desirable to fabricate housing 70 from Ti--Nb--Zr toreduce the risk of thrombogenesis.

FIGS. 8A and B are schematic diagrams of the Texas HeartInstitute/Abiomed TAH design which utilizes a d.c. motor to drive aminiature centrifugal pump 80 that rotates at about 6,000-8,000 RPM.This pump 80 pressurizes hydraulic fluid alternately into chambers 82and 84 separated by septum 86 and enclosed by flexible diaphragms 88 and90 respectively. As fluid is pumped into chamber 82, diaphragm 88expands into heart space 92 forcing blood from this space. At the sametime, fluid is pumped from chamber 84 causing diaphragm 90 to relax andexpanding heart space 94, drawing blood into the TAH. The hydraulic flowis reversed by a two-position 4-way rotating valve 100 of radialconfiguration for compactness. Rotary valve 100 rotates within sleeves102 and 104 at high speed so that contacting-surfaces between the valve100 and these sleeves are subject to wear. Further, rotary valve 100rotates against seals 106 and wear may be expected at the contactingsurfaces of the seals and the valve 100.

Several components of the Texas Heart Institute/Abiomed TAH may befabricated according to the invention. Thus, high speed components ofthe centrifugal pump 80 subject to wear may be fabricated fromTi--Nb--Zr alloy and then surface hardened and/or coated with a tightlyadherent, hard, wear resistant coating. Further, the rotary valve 100itself and the surfaces of sleeves 102, 104 and seals 106 may befabricated from Ti--Nb--Zr alloy then surface hardened or coated with anadherent, wear-resistant coating. Finally, the inner surfaces of the TAH108, 110 may be fabricated from Ti--Nb--Zr alloys to improve bloodcompatibility and reduce the potential for thrombus and blood clotting.

The Novacor designed VAD illustrated in FIGS. 9A and B have a solenoidmechanism 120 which sends energy through beam-springs 122, 124 thatextend to the back of pump pusher plates 126, 128. The energy stored inthe springs translates into linear motion of the plates which exertsforce on the flexible blood sac 130. The blood sac 130 consists of abutyl rubber layer sandwiched between two layers of polyurethane Biomer.The blood sac 130 is supported within a cylindrical aluminum ring 132that acts as a pump housing. The blood inflow 133 and outflow 134 portsare positioned tangentially on opposite sides of the housing to ensurestraight-through blood flow. The ports are formed of anepoxy-impregnated Kevlar fabric shell that-is integrated into thehousing. The ports also encapsulate trileaflet inlet and outlet valvesmade from bovine pericardium tissue. When implanted into the body,fittings for attaching inflow and outflow valves to vascular conduitsare bonded to a pump bulkhead, not shown, which also provides theframework for an encapsulating shell around the pump. This encapsulatingshell also has provision for mounting the solenoid energy converter. Thesolenoid energy converter consists of two solenoid mechanisms, twolightweight titanium beam-springs, and an aluminum support structure.All of these metallic components would come into contact with bloodcomponents and body tissue. Therefore, the invention proposes that thetitanium beam-springs be replaced with beam-springs of Ti--Nb--Zr alloy.Further, the aluminum support structure would likewise be replaced witha Ti--Nb--Zr alloy support structure that may optionally be hardenedand/or coated with a hard coating.

Novacor has identified, in designing the solenoid, that "the challengewas coming up with something that would run for 100 million cycles ayear, without requiring maintenance." O'Connor, Lee, "Novacor's VAD: Howto Mend a Broken Heart," Mechan. Engr'g pp. 53-55 (November 1991). Theinvention components fabricated from Ti--Nb--Zr alloys then hardened orcoated with hard, wear resistant coatings provide surfaces that arehard, microfretting wear resistant, biocompatible and blood compatibleso that they would meet this goal. To further reduce friction and wearof wear surfaces of implant devices, a thin boron or silver surfacelayer can be applied as an overlay on the previously diffusion hardenedTi--Nb--Zr surface.

External mechanical hearts (EMHs) are used as a bridge to transplant.These hearts include the Jarvik-7 pneumatic heart and the more recentleft-ventricular assist device, the Heartmate developed by ThermocardioSystems. In the Heartmate system, two tubes, one carrying air and theother electrical wire, pass from outside the body to an implanted bloodpump. The pump is implanted in the abdomen and removes blood from thenatural heart's left ventricle. This blood enters and exits the pumpthrough 25 millimeter input and output valves made from chemicallyprocessed bovine tissue. The blood flows from the output valve through adacron-wrapped polyurethane tube to the aorta. An electric motor mountedin the Heartmate's lower chamber actuates a flat-plate piston, which isbonded to a flexible polyurethane diaphragm. When the motor goes throughone revolution, it turns a cam assembly that compresses the diaphragm,which pushes blood through the output valve. The operation of the pumpis controlled by a microprocessor located in a shoulder bag whichadjusts the heartbeat rate by changing the motor's current. According tothe invention, the moving parts of the heartmate pump may be replacedwith components fabricated from Ti--Nb--Zr alloys then hardened orcoated with a hard coating to reduce mechanical wear, friction, andmicrofretting wear. Furthermore, those metallic components that comeinto contact with blood components, may also be replaced with Ti--Nb--Zralloy components similarly coated to improve blood compatibility andreduce the risk of clot-formation.

The gravest problem in the use of the pneumatic Jarvik-7 heart has beenidentified as the formation of blood clots. O'Connor, Lee, "Engineeringa Replacement for the Human Heart," Mechan. Engr'g pp. 36-43 (July1991). In 1990, the FDA withdrew the Jarvik system from clinical trialsdue to concerns over quality control during manufacture. The Universityof Utah made modifications to the design of the Jarvik heart to developa new system called the "Utah 100" which has elliptical pump housings,as opposed to the spherical housings of the Jarvik-7. Further, the Utah100 has redesigned Junctions for joining the diaphragms within theventricles to the housing. These changes are said to have resulted in anabout 70% reduction in blood clot formation relative to the Jarvik-7design. However, according to the invention, yet further reduction inblood clot formation may be obtained by fabricating moving parts andthose metallic surfaces that contact blood components from Ti--Nb--Zralloys and then hardening and/or coating these components with hard,wear-resistant coating to increase blood compatibility and thrombusresistance, and to reduce abrasive wear, and reduce microfretting wear.

Guide Wires and Catheters

FIGS. 13A and B show, in partial cross section, the distal end of aguide wire fabricated according to the invention. The guide wire 145 hasa core 140 of Ti--Nb--Zr alloy with a surface hardened coating 142 toreduce friction which may be further coated with a material that is bio-and hemocompatible and of low friction when in contact with the catheterwall or body tissue. The flexible catheter 146 through which the guidewire moves is also fabricated according to the invention and includes acoil 147 of low modulus titanium alloy which is generally encased by apolymer sheath 148 as shown schematically in FIG. 13C. The guide wire inthis case is equipped with a cutting tip 144, preferably also made ofTi--Nb--Zr alloy with a diffusion hardened surface optionally with ahard ceramic or lubricating coating. Since the guide wire is fabricatedof metal, it is highly visible under X-rays, providing excellentradiopacity. Boron or silver surface layers may also be deposited on thediffusion hardened surfaces to further reduce friction and wear.

Pacemakers and Electrical Signal Carrying Leads/Sensors

Pacemaker and other electronic leads are manufactured by severalcorporations, including Medtronic, which produces a range of pacemakerlead designs.

One of these designs is shown in schematic form in FIGS. 14A and B. Thepacemaker lead body 150 has a centrally disposed metallic conductor 152typically made of cobalt-nickel alloy, such as MP35N®. This conductor152 is usually made up of several strands of wire, each having adiameter of about 0.15-0.20 mm. The conductor 152 is covered by aninsulative, protective polymer sheath 153 so that the elongate body 150of the pacemaker lead has an overall diameter ranging from about 2.2 toabout 3 mm. The pacemaker has a first end 154 with an electrode 158 forconnecting to a pulse generator and a second end 156 with an electrode157 for contacting heart muscle. An alternative embodiment is shown inFIG. 14C. As supplied, these two ends are covered with protectivepolyurethane caps which can be removed for installation of thepacemaker. In order to prevent electrical interference with theconductor 152, a polymeric insulative sleeve 153 is disposed over theentire pacemaker lead body 150, with the exception of the exposedelectrodes 157 for contacting heart muscle and the contact electrode 158for engaging with the pulse generator that houses the electronics andpower pack for the pacemaker. As explained before, the organic polymericsheath compositions, typically polyurethane, can slowly degenerate inthe body causing problems, not only due to potential deterioration ofelectrical insulation and interference with electrical signals but alsobecause of potentially toxic products of degradation.

The invention provides, as shown in FIG. 14C, a pacemaker wherein theconductor 152 is fabricated from a Ti--Nb--Zr alloy that is coated witha tightly adherent, low friction, bio- and hemocompatible coating, withthe exception of the electrode for contacting heart muscle 157, and theelectrode 158 at the other end of the lead for engaging the pulsegenerator. The coatings can be formed by in situ oxidation or nitridingof the Ti--Nb--Zr to produce an electrically insulative surface layer offrom about 0.1 to about 3 microns in thickness, preferably less thanabout 0.5 microns in thickness. This process can be carried out at thesame time the material is age-hardened. Alternatively, an insulativeinert ceramic coating can be applied by conventional CVD or PVD methodseither on the original Ti--Nb--Zr alloy surface or onto the diffusionhardened Ti--Nb--Zr surface. For these overlay coatings, the thicknesscan be as great as 20 microns. The overlay coatings include ceramicmetal oxides, metal nitrides, metal carbides, amorphous diamond-likecarbon, as detailed above. The electrical signal conductor 152 cancomprise either a single wire or multiple wires. Exposed Ti--Nb--Zrmetallic ends of the wire or wires are preferably connected directly toa pulse generator thereby avoiding the necessity for a weld or crimp toattach an electrode to the conductor which may result in local galvaniccorrosion or physically weakened regions. Further, since the coatingsprovide a natural protective insulative surface, the use of a coiledconstruct could be avoided by using only a preferred single-strand,non-coiled low modulus Ti--Nb--Zr metallic wire construct for theconductor 152. This will also eliminate the need for stiff guide wire.Finally, the overall diameter of the pacemaker lead body 150 could bereduced considerably from the range of about 2.2-3 mm for currentcommercially available leads to about 0.2-1 mm. Optionally, the leads ofthe invention may be covered with a polymeric sheath.

Stents

FIG. 11A shows a schematic of an expandable stent 160, in non-expandedstate, positioned on the distal end of a balloon expandable segment 162of a guide wire 164. The stent is fabricated from Ti--Nb--Zr alloy andis designed so that it can be collapsed over a balloon segment of aballoon catheter. When the stent is in position, within segment of atubular conduit 165 in the body, a blood vessel for example, to bepropped open, the balloon 162 is expanded thereby expanding the stent160 radially outward up to the blood vessel wall 166 so that means forgripping soft tissue, such as barbs (not shown), on the outer surface ofthe stent 160 engage and grip blood vessel tissue to anchor the stent160 in position as shown in FIG. 11B. The balloon 162 is then collapsedand removed leaving the stent in place as shown in FIG. 11C. In thisway, the blood vessel is permanently propped open. Urinary,gastrointestinal, and other stent applications are also provided usingTi--Nb--Zr alloy.

Grafts

FIG. 10 is a representative sketch of a side view of a substantiallytubular vascular graft 170 sized to graft onto a blood vessel and madeof woven low modulus Ti--Nb--Zr wires 172. While the graft shown is madeof woven wires of Ti--Nb--Zr, the graft can also be fabricated from acylindrical tubing of this alloy. The graft can be fabricated fromTi--Nb--Zr alloy in the lower modulus cold worked condition, or in theslightly higher modulus aged condition with optional surface hardening.Additionally, protein, antibiotic, anti-thrombic, and other surfacetreatments may be employed to further improve the biocompatibility andclinical performance.

Defibrillators

FIGS. 12A and B show a defibrillator including a flexible siliconepolymeric patch 300 with a coil of conductive wire 320 (typicallytitanium, stainless steel, or cobalt-nickel-chromium) on the side of thesilicone patch 300 that will contact muscle tissue. When in place in thebody, the lead wire 320 that carries power to the coil 340 extends outof the body (through the skin) and is electrically connected to a powersource contained in a protective container 360. According to theinvention, the lead wire 320 is fabricated with an electricallyconductive core 350 of Ti--Nb--Zr alloy and is coated with an adherentelectrically insulative coating 380, such as metal oxides, carbides, ornitrides, or with amorphous diamond-like carbon as shown in exaggerateddetail FIG. 12B. This coating electrically insulates the lead wire fromelectrical contact with surrounding body tissue while also protectingthe metallic core from corrosion and attack by body fluids, as describedpreviously for the pacemaker lead. Elimination oft he polymer coatingresults in the elimination of potentially toxic products of gradualdegradation of the polymer and also the consequent shorting the systemwhen the insulative coating is breached.

The Hardened Surfaces

The oxygen or nitrogen diffusion hardened surface of the alloy implantsmay be highly polished to a mirror finish to further improve blood flowcharacteristics. Further, the oxide- or nitride-coated surfaces maybecoated with substances that enhance biocompatibility and performance.For example, a coating of phosphatidyl choline, heparin, or otherproteins to reduce platelet adhesion to the surfaces of the implant, orthe use of antibiotic coatings to minimize the potential for infection.Boronated or silver-doped hardened surface layers on the implant reducesfriction and wear between contacting parts of heart valves, prostheticartificial hearts, ventricular assist devices, and other contactingparts in the invention cardiovascular implants. Additionally, amorphousdiamond-like carbon, pyrolytic carbon, or other hard ceramic surfacelayers can also be coated onto the diffusion hardened surface tooptimize other friction and wear aspects. The preferred diffusionhardened surface layer described in this application provides a hard,well-attached layer to which these additional hard coatings can beapplied with a closer match between substrate and coating with respectto hardness. Other, conventional methods of oxygen surface hardening arealso useful. Nitriding of the substrate leads to a hardened nitridesurface layer. Methods of nitridation known in the art may be used toachieve a hard nitride layer.

Regardless of how a Ti--Nb--Zr alloy implant's surface is hardened, thefriction and wear (tribiological) aspects of the surface can be furtherimproved by employing the use of silver doping or boronation techniques.Ion-beam-assisted deposition of silver films onto ceramic surfaces canimprove tribiological behavior. The deposition of up to about 3 micronsthick silver films can be performed at room temperature in a vacuumchamber equipped with an electron-beam hard silver evaporation source. Amixture of argon and oxygen gas is fed through the ion source to createan ion flux. One set of acceptable silver deposition parameters consistsof an acceleration voltage of 1 kev with an ion current density of 25microamps per cm². The silver film can be completely deposited by thision bombardment or formed partially via bombardment while the remainingthickness is achieved by vacuum evaporation. Ion bombardment improvesthe attachment of the silver film to the Ti--Nb--Zr alloy substrate.Similar deposition of silver films on existing metal cardiovascularimplants may also be performed to improve tribiological behavior, aswell as antibacterial response.

An alternate method to further improve the tribiological behavior ofTi--Nb--Zr alloy surfaces of cardiovascular implants is to applyboronation treatments to these surfaces such as commercial availableboride vapor deposition, boron ion implantation or sputter depositionusing standard ion implantation and evaporation methods, or form aboron-type coating spontaneously in air. Boric Acid (H₃ BO₃) surfacefilms provide a self-replenishing solid lubricant which can furtherreduce the friction and wear of the ceramic substrate. These films formfrom the reaction of the B₂ O₃ surface (deposited by variousconventional methods) on the metal surface with water in the body toform lubricous boric acid. Conventional methods that can be used todeposit either a boron (B) , H₃ BO₃, or B₂ O₃ surface layer on thecardiovascular implant surface include vacuum evaporation (with orwithout ion bombardment) or simple oven curing of a thin layer over theimplant surface. The self-lubricating mechanism of H₃ BO₃ is governed byits unique layered,-triclinic crystal structure which allows sheets ofatoms to easily slide over each other during articulation, thusminimizing substrate wear and friction.

Additionally, surfaces (metal or coated) of all the cardiovascular andmedical implants discussed may optionally be coated with agents tofurther improve biological response. These agents includeanticoagulants, proteins, antimicrobial agents, antibiotics, and thelike medicaments.

Although the invention has been described with reference to itspreferred embodiments, those of ordinary skill in the art may, uponreading this disclosure, appreciate changes and modifications which maybe made and which do not depart from the scope and spirit of theinvention as described above and claimed below.

What is claimed is:
 1. A woven vascular graft for grafting onto a bloodvessel, the woven vascular graft having enhanced hemocompatibility,comprising:a woven vascular graft at least partially fabricated from ametal alloy comprising:(i) titanium; and (ii) from about a range of 10to about 20 wt. % niobium or from about 35 to about 50 wt. % niobium;wherein the alloy is free of deliberately added toxic elements, exceptsuch amounts of said toxic elements as may occur as impurities in thealloy and as contaminants as a result of an alloying process.
 2. Thewoven vascular graft of claim 1, wherein said alloy further comprises anamount of zirconium from about 0.5 to about 20 wt. % zirconium which issufficient to retard the transformation of β-phase titanium.
 3. Thewoven vascular graft of claim 2, wherein the metal alloy comprises fromabout 0.5 to about 20 wt. % zirconium.
 4. The woven vascular graft ofclaim 1, further comprising a hard outer surface layer on at least aportion of the woven vascular graft, said layer formed on the wovenvascular graft by a process selected from the group consisting of oxygendiffusion hardening, nitrogen diffusion hardening, physical vapordeposition, and chemical vapor deposition.
 5. The woven vascular graftof claim 2, further comprising a hardened outer surface layer on atleast a portion of the woven vascular graft, said layer formed on thewoven vascular graft by a process selected from the group consisting ofoxygen diffusion hardening, nitrogen diffusion hardening, physical vapordeposition, and chemical vapor deposition.
 6. The woven vascular graftof claim 1 or claim 2 or claim 3 or claim 4, wherein surfaces of thewoven vascular graft that come into contact with body tissue or fluidare at least partially coated with a composition selected from the groupconsisting of anticoagulants, antimicrobial agents, antibiotics, andmedicaments.
 7. The woven vascular graft of claim 1 or claim 2 or claim3 or claim 4, further comprising a lower friction wear-resistant outersurface layer on at least a portion of the woven vascular graft, saidlower friction wear-resistant outer surface layer produced by a processselected from the group consisting of boronation and silver doping. 8.The woven vascular graft of claim 2, wherein the metal alloy comprisesabout 74 wt. % titanium, about 13 wt. % niobium, and about 13 wt. %zirconium.
 9. The woven vascular graft of claim 2, wherein the metalalloy comprises titanium, from about 10 to about 20 wt. % niobium, andup to about 20 wt. % zirconium.
 10. The woven vascular graft of claim 2,wherein the metal alloy comprises titanium, from about 35 to about 50wt. % niobium, and up to about 20 wt. % zirconium.
 11. A woven vasculargraft for grafting onto a blood vessel, the woven vascular graft havingenhanced hemocompatibility, comprising:a woven vascular graft at leastpartially fabricated from a metal alloy comprising:(i) titanium; and(ii) niobium and tantalum, wherein the combined wt. % of niobium andtantalum is from a range of about 10 to about 20 wt. % or from about 35to about 50 wt. %; wherein the alloy is free of deliberately added toxicelements, except such amounts of said toxic elements as may occur asimpurities in the alloy and as contaminants as a result of an alloyingprocess.
 12. The woven vascular graft of claim 11, wherein said alloyfurther comprises an amount of zirconium from about 0.5 to about 20 wt.% zirconium which is sufficient to retard the transformation of β-phasetitanium.
 13. The woven vascular graft of claim 12, wherein the metalalloy comprises from about 0.5 to about 20 wt. % zirconium.
 14. Thewoven vascular graft of claim 11, further comprising a hardened outersurface layer on at least a portion of the woven vascular graft, saidlayer formed on the woven vascular graft by a process selected from thegroup consisting of oxygen diffusion hardening, nitrogen diffusionhardening, physical vapor deposition, and chemical vapor deposition. 15.The woven vascular graft of claim 12, further comprising a hard outersurface layer on at least a portion of the woven vascular graft, saidlayer formed on the woven vascular graft by a process selected from thegroup consisting of oxygen diffusion hardening, nitrogen diffusionhardening, physical vapor deposition, and chemical vapor deposition. 16.The woven vascular graft of claim 11 or claim 12 or claim 13 or claim14, wherein surfaces of the woven vascular graft that come into contactwith body tissue or fluid are at least partially coated with acomposition selected from the group consisting of anticoagulants,antimicrobial agents, antibiotics, and medicaments.
 17. The wovenvascular graft of claim 11 or claim 12 or claim 13 or claim 14, furthercomprising a low friction wear-resistant outer surface layer on at leasta portion of the woven vascular graft, said lower frictionwear-resistant outer surface layer produced by a process selected fromthe group consisting of boronation and silver doping.